Implantable micro-component electrodes

ABSTRACT

The present disclosure provides robust implantable micro-component electrodes that can be used in a variety of medical devices. The medical device may be a neural probe that can monitor or stimulate neural activity in an organism&#39;s brain, spine, nerves, or organs, for example. The micro-component electrode has a small physical profile, with ultra-thin dimensions, while having high strength and flexibility. The micro-electrode has an electrically conductive core material, e.g., carbon. The surface of the core material includes one or more electrically conductive regions coated with an electrically conductive material and one or more non-conductive regions having an electrically non-conductive biocompatible polymeric coating. Implantable devices having such micro-components are capable of implantation in an organism for very long durations.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a 371 U.S. National Stage of InternationalApplication No. PCT/US2011/040741, filed Jun. 16 , 2011, which claimsthe benefit of U.S. Provisional Application No. 61/356,482, filed onJun. 18, 2010. The entire disclosures of the above applications areincorporated herein by reference.

GOVERNMENT RIGHTS

This invention is made with government support under NS068396 andEB002030 awarded by the National Institutes of Health. The governmenthas certain rights in the invention.

FIELD

The present disclosure relates to conductive micro-components, and morespecifically to conductive micro-component electrodes capable ofimplantation in an organism to monitor electrophysiological and/orelectrochemical activity, such as a neural probe for monitoring neuralactivity.

BACKGROUND

This section provides background information related to the presentdisclosure which is not necessarily prior art.

Implantable microscale neural probes are important tools forneuroscience. The capability of monitoring specific neuronal ensemblesfor long periods of time with great precision is a powerful tool inneuroscience research for linking low-level neuronal circuits tohigh-level brain function, such as learning, memory, and perception. Onthe clinical side, it also enables the development of closed-loopneurostimulation and neuroprosthetic systems using detailedneurophysiological signals for feedback. Beyond neural recordingapplications, such microelectrodes can enable new approaches forcreating long-lasting, high-fidelity neural interfaces that woulddirectly benefit neurostimulation applications.

Since the pioneering work of Strumisser in 1958 introducing microwirebundles for chronic neural recordings in hibernating squirrels, therehas been an ongoing push to develop improved implantable microelectrodetechnologies. While various implantable probes have been investigated,the recording quality is not optimally high. Further, eventually therecordings degrade and ultimately fail over time with conventionalimplantable microelectrodes. The primary challenge in neural interfacetechnologies today remains that of making reliable implantable devicesfor long-term, stable, high-fidelity spike recordings from selectiveneuronal ensembles.

SUMMARY

This section provides a general summary of the disclosure, and is not acomprehensive disclosure of its full scope or all of its features.

In various aspects, the present disclosure provides implantablemicro-component electrodes. In certain aspects, the present disclosurepertains to an implantable micro-component electrode that comprises anelectrically conductive core material. In certain aspects, theconductive core material may have an elastic tensile modulus of greaterthan or equal to about 200 GPa and a nominal value of stiffness ofgreater than or equal to about 5 GN/μm. In certain variations, theelectrically conductive core material comprises carbon. Further, theimplantable micro-component electrode comprises one or more electricallyconductive regions disposed on the surface of the electricallyconductive core material. The one or more electrically conductiveregions comprise an electrically conductive biocompatible coating.Further, the implantable micro-component electrode comprises anelectrically non-conductive biocompatible coating disposed on regions ofthe surface of the electrically conductive core material. The regionswhere the electrically non-conductive biocompatible coating is disposedare those where the one or more electrically conductive regions areabsent. In various aspects, the micro-component electrode has at leastone dimension of less than or equal to about 10 μm. In certain aspects,the electrically non-conductive biocompatible coating has a thickness ofless than 1 μm.

In other aspects, the present teachings provide a micro-componentelectrode for an implantable medical device that comprises anelectrically conductive core material comprising carbon and having adiameter of less than or equal to about 10 μm. One or more electricallyconductive regions are disposed on the surface of the electricallyconductive core material. The one or more electrically conductiveregions comprise an electrically conductive polymeric coating. Further,an electrically non-conductive polymeric coating is disposed on regionsof the surface of the electrically conductive core material where theone or more electrically conductive regions are absent. In certainvariations, the electrically non-conductive coating comprises a parylenepolymer or a parylene polymer derivative. In certain variations, theelectrically conductive non-conductive coating has a thickness of lessthan or equal to about 1 μm.

In yet other aspects, the present disclosure provides methods ofmonitoring, sensing, or stimulating neural activity in an organism. Inone aspect, such a method comprises electrically communicating with aprobe implanted into an organism. In certain variations, the methodcomprises electrically communicating with a neural probe implanted intoa brain of an organism. The probe optionally has a cross-sectional areaof less than or equal to about 2,000 micrometers-squared (μm²) andcomprises at least one micro-component electrode. The at least onemicro-component electrode comprises an electrically conductive corematerial having a surface with one or more electrically conductiveregions disposed on the surface of the electrically conductive corematerial. The one or more electrically conductive regions disposed onthe surface of the electrically conductive core comprise an electricallyconductive polymeric coating. Further, an electrically non-conductivepolymeric coating is disposed on regions of the surface of theelectrically conductive core material corresponding to locations wherethe one or more electrically conductive regions are absent. In certainvariations, the micro-component electrode has at least one dimensionless than or equal to about 10 μm.

Further areas of applicability will become apparent from the descriptionprovided herein. The description and specific examples in this summaryare intended for purposes of illustration only and are not intended tolimit the scope of the present disclosure.

DRAWINGS

The drawings described herein are for illustrative purposes only ofselected embodiments and not all possible implementations, and are notintended to limit the scope of the present disclosure.

FIGS. 1A-1E show comparative performance of micro-component electrodes,some of which are prepared in accordance with the present teachings.FIGS. 1A-1C are scanning electron microscopy SEM images of comparativebare carbon fibers (1A), 800 nm parylene-N coated carbon fibers (1B),and parylene coated carbon fiber with an electrodeposited PEDOT/PSSelectrode recording site (1C). FIG. 1D is an assembly of micro-componentmicrothread electrodes (MTE); carbon fibers are mounted onto a printcircuit board, chemical vapor deposition (CVD) parylene coated,functionalized parylene coated with poly(ethylene glycol) (PPX-PEG)coated, cut, and Poly(3,4-ethylenedioxythiophene) (PEDOT)electrodeposited. FIG. 1E is a size comparison of a 5 mm “Michigan”style electrode and a Microthread Electrode prepared with themicro-component electrodes of the present disclosure.

FIGS. 2A-2E show in vitro characterization of Microthread Electrodes.FIG. 2A is Cyclic Voltammetry traces (CV) of parylene coated fiber,parylene coated carbon fiber with an exposed carbon tip, and parylenecoated carbon fiber with an electrodeposited PEDOT/PSS recording siteusing various deposition charge densities. FIG. 2B is a Bode magnitudeimpedance plot. FIG. 2C is a Bode phase plot. FIG. 2D is a comparison at1 kHz impedance. And FIG. 1E is a comparison of charge carryingcapacity.

FIGS. 3A-3E show in vivo single unit recording capabilities ofmicro-component electrodes prepared in accordance with the presentteachings. FIG. 3A is a piled single unit neural recordings from aparylene coated PEDOT microthread electrode. FIG. 3B is an average unitspike. FIGS. 3C and 3D are five seconds of raw neural recordings from aPEDOT site MTE (3C) and a comparative carbon site MTE (3D). FIG. 3Eshows five seconds of Local Field Potential (LFP) recordings from aPEDOT site MTE and a comparative carbon site MTE (3E).

FIGS. 4A-4C show in vivo characterization of micro-component electrodesprepared in accordance with the present teachings. FIG. 4A show a PowerSpectral Density plot of carbon site MTE and PEDOT site MTE (4A). FIGS.4B and 4C show Spectrograms of comparative PEDOT site MTE (4B) andcarbon site MTE (4C).

FIGS. 5A-5B′ show a neural probe design with sub-cellular dimensionsusing a thin lateral platform. FIG. 5A is an SEM perspective view of aparylene-based open architecture device used for in vivo testing. Thetip of the device is at the lower left. FIG. 5B is a CAD drawing ofprobe design indicating overall length and width of the latticestructures (4 μm), and FIG. 5B′ is a cross-sectional view of line A-A′shown in (FIG. 5B). Scale of bars is 100 μm.

FIGS. 6A-6B are an illustration of the biomimetic principle ofsub-cellular size to modulate a foreign body response. FIG. 6A has twostructures having a similar tissue response after 4-weeks, but nothaving sub-cellular dimensions and FIG. 6B is a cross-section of a SEEprobe. Each of FIGS. 6A-6C shows microglia, which are large relative tothe 5 μm edge.

FIGS. 7A-7D show examples of qualitative and quantitative results arounda non-functional probe. FIG. 7A shows GFAP and OX42 antibodies labelastrocytes and microglia, respectively. FIG. 7B shows GFAP and NeuNlabeling astrocytes and neuronal nuclei. FIG. 7C is normalized meannonneuronal density as a functional of distance from probe interface andFIG. 7D is a mean neuronal density. P<0.05 and scale is 100 μm.

FIGS. 8A-8C shows diagrams of assemblies of micro-components prepared inaccordance with the present disclosure. FIG. 8A shows a single-strandmicro-component electrode. FIG. 8B shows a side profile sectional viewof the single-strand micro-component electrode. FIG. 8B shows amulti-strand microthread probe illustrating the flexible carbon nanotube(CNT) composite core and electrode site with a thin, conformal coatingof an insulating and functionalized polymer. The CNT is nominally 5 μmon a side and has a 0.5 to 1 μm thick coating to give the probe asub-cellular dimension.

FIG. 9 is a schematic of the layer-by-layer assembly process that can beused to form core composite materials for the micro-components of thepresent disclosure. The inset shows an SEM micrograph of cultured neuronon SWNT multilayers.

FIG. 10 shows CVD polymerization of various [2.2]paracyclophanes tocreate functionalized non-conductive parylene coatings on substrates foruse in conjunction with the present teachings. Corresponding referencenumerals indicate corresponding parts throughout the several views ofthe drawings. The lower portion of the drawing shows spatiallycontrolled CVD deposition.

FIGS. 11A-11F shows an exemplary formation of parylene coating andfunctionalizing polymeric coatings by atom-transfer radicalpolymerization (ATRP) applied to a conductive core material to form amicro-component of the present disclosure. In FIG. 11A carbon fibers arecoated with 800 nm poly(p-xylylene). In FIG. 11B, the fiber of FIG. 11Ais further coated with a 50 nm layer ofpoly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)]. In FIG.11C, poly(ethylene glycol) (PEG) is covalently grafted ontopoly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)] throughATRP. FIG. 11D shows removal of insulation by cutting away the tip toexpose a carbon site, and then electrodeposited with PEDOT. FIGS.11E-11F are SEM images of such modified MTEs prepared in accordance withthe present teachings.

FIGS. 12A-12D show exemplary schematics and cross section ofarchitecture of a neural probe incorporating a micro-component electrodein accordance with one aspect of the present teachings. FIG. 12A shows afirst side view of an exemplary neural probe. FIG. 12B shows a detailedsectional view taken along line B-B in FIG. 12A. FIG. 12C shows a secondside view of another exemplary neural probe. FIG. 12D shows a detailedsectional view of the layers forming an exemplary MTE probe.

FIGS. 13A-13F. FIG. 13A shows scanning electron microscopy (SEM) imagesof comparative bare carbon fibers, 800 nm parylene-N coated carbon fiberwith an exposed conductive carbon tip site, and parylene coated carbonfiber with an electrodeposited platinum black recording site. FIG. 13Bshow comparative charge storage capacity. FIG. 13C shows comparative 1kHz impedance. FIG. 13D is cyclic voltammetry traces (CV) of bare carbonfibers, 800 nm parylene-N coated carbon fiber with an exposed conductivecarbon tip site, and parylene coated carbon fiber with anelectrodeposited platinum black recording site. FIG. 13E showscomparative electrochemical sensitivity to hydrogen peroxide duringconstant potential amperometry. FIG. 13F shows comparative performancesof electrical recording capabilities.

FIGS. 14A-14L show long-term neural electrophysiological recordingcapability in the brain. FIG. 14A shows percent of active chronicallyimplanted MTEs able to detect at least 1 single unit (dashed line) as afunction of weeks post-implant (n=7). FIG. 14B shows mean SNR of thelargest single unit detected on each electrode. FIG. 14C shows meanamplitude of largest single unit detected on each electrode (solid,black), and the mean noise floor of each electrode (dashed, red). FIG.14 has amplitude of single units from longest implant (solid, black).Amplitude of noise floor from same animal (dashed, red). FIGS. 14E-Hshow electrophysiological recordings taken from a rat with a MTEimplanted in M1 five weeks post-implant. FIGS. 14I-L showelectrophysiological recordings taken from a different rat implantedwith a MTE in M1 seven weeks post-implant. FIGS. 14E and 14I shows meanwaveform of discernable single units. FIGS. 14F and 14J show piledsingle units from two minutes of recordings. FIGS. 14G and 14K showrepresentative example of two seconds of high-speed recordings. FIGS.14H and 14L show results from principal component analysis showingdistinct clusters.

DETAILED DESCRIPTION

Example embodiments are provided so that this disclosure is thorough,and will fully convey the scope to those who are skilled in the art.Numerous specific details are set forth such as examples of specificcomponents, devices, and methods, to provide a thorough understanding ofembodiments of the present disclosure. It is apparent to those skilledin the art that specific details need not be employed, that exampleembodiments may be embodied in many different forms and that neithershould be construed to limit the scope of the disclosure. In someexample embodiments, well-known processes, well-known device structures,and well-known technologies are not described in detail.

The terminology used herein is for the purpose of describing particularexample embodiments only and is not intended to be limiting. As usedherein, the singular forms “a,” “an,” and “the” may be intended toinclude the plural forms as well, unless the context clearly indicatesotherwise. The terms “comprises,” “comprising,” “including,” and“having,” are inclusive and therefore specify the presence of statedfeatures, integers, steps, operations, elements, and/or components, butdo not preclude the presence or addition of one or more other features,integers, steps, operations, elements, components, and/or groupsthereof. The method steps, processes, and operations described hereinare not to be construed as necessarily requiring their performance inthe particular order discussed or illustrated, unless specificallyidentified as an order of performance. It is also to be understood thatadditional or alternative steps may be employed.

When an element or layer is referred to as being “on,” “engaged to,”“connected to,” or “coupled to” another element or layer, it may bedirectly on, engaged, connected or coupled to the other element orlayer, or intervening elements or layers may be present. In contrast,when an element is referred to as being “directly on,” “directly engagedto,” “directly connected to,” or “directly coupled to” another elementor layer, there may be no intervening elements or layers present. Otherwords used to describe the relationship between elements should beinterpreted in a like fashion (e.g., “between” versus “directlybetween,” “adjacent” versus “directly adjacent,” etc.). As used herein,the term “and/or” includes any and all combinations of one or more ofthe associated listed items.

Although the terms first, second, third, etc. may be used herein todescribe various elements, components, regions, layers and/or sections,these elements, components, regions, layers and/or sections should notbe limited by these terms. These terms may be only used to distinguishone element, component, region, layer or section from another region,layer or section. Terms such as “first,” “second,” and other numericalterms when used herein do not imply a sequence or order unless clearlyindicated by the context. Thus, a first element, component, region,layer or section discussed below could be termed a second element,component, region, layer or section without departing from the teachingsof the example embodiments.

Spatially relative terms, such as “inner,” “outer,” “beneath,” “below,”“lower,” “above,” “upper,” and the like, may be used herein for ease ofdescription to describe one element or feature's relationship to anotherelement(s) or feature(s) as illustrated in the figures. Spatiallyrelative terms may be intended to encompass different orientations ofthe device in use or operation in addition to the orientation depictedin the figures. For example, if the device in the figures is turnedover, elements described as “below” or “beneath” other elements orfeatures would then be oriented “above” the other elements or features.Thus, the example term “below” can encompass both an orientation ofabove and below. The device may be otherwise oriented (rotated 90degrees or at other orientations) and the spatially relative descriptorsused herein interpreted accordingly.

Example embodiments will now be described more fully with reference tothe accompanying drawings.

Implantable neural interface devices are important to a broad class ofemerging neuroprosthetic and neurostimulation systems, in both researchand clinical settings. In almost all cases, the performance of thesystem hinges to a large degree on the performance of the device torecord and/or stimulate within quality, stability, and longevityrequirements. Recording quality, longevity and stability is highlyvariable and the reactive tissue response that occurs to devicesfollowing implantation is a likely key contributing factor todiminishing the device performance. The fundamental challenge is todevelop advanced materials and implantable structures that will enableneural interface devices to be implanted in vivo in target areas of anorganism, such as in a brain, and remain functional for long durations,potentially for several years to several decades.

Penetrating neural probe technologies have allowed investigators torecord both chemical and electrical signals in the brain. However,implantation of these neural probes causes acute tissue trauma withneuronal injury, recruitment of microglia, activation of astrocytes, andmost importantly, disruption of the blood brain barrier (BBB) followedby plasma protein adsorption onto the surface of the probe. This acutetrauma can cause abnormal electrophysiological responses, temporaryincreases in neurotransmitter levels, affect coating technologies, andperpetuate chronic immune responses.

While conventional carbon fiber microelectrodes can record single unitactivity, patch clamp neural activity, and changes in extracellulardopamine concentrations, their ability to perform in vivo as a chronicrecording device has been limited by the site size necessary to obtainsingle unit recordings and the insulation material, such as glass orfused silica, which has the negative effect of increasing the device'sfootprint and stiffness. Further computer models and experimentalstudies of the probe-tissue interface suggest that probes having greaterflexibility may help to minimize perpetual mechanical trauma cause byphysiological motion between the probe and surrounding tissue, ascompared to trauma caused by conventional probe devices.

Recently, there has been an increased understanding of the detailedreactive tissue responses to implantable probes. Tissue encapsulation ofthe probe coupled with neuronal death in the vicinity of the recordingelectrode, have been implicated as the two largest variables negativelyimpacting the stability and longevity of long-term neural recordings.Regardless of the microelectrode technology used, cellular and acellularencapsulation forms around the electrode site. Tissue encapsulation hasbeen implicated in an increase in electrode impedance, a decline insignal amplitude, increased noise, and possibly even the silencing ofneurons. Many different approaches have failed to improve recordingstability or mitigate the encapsulation around neural devices, includingelectrode surface modification, and stem cell delivery. Further,reduction of tissue encapsulation has also proven elusive for variousconventional implantable, microscale sensors.

Developing smaller and more flexible neural probes with improved surfacechemistry for long-term, high quality and selective neural recording isimportant for both neuroscience research and clinical neurotechnologies.Neural probes formed in accordance with the principles of the presentdisclosure, provide can provide monitoring of specific neuronalensembles for long periods of time with great precision, which is apowerful tool in neuroscience research for linking low-level neuronalcircuits to high-level brain function, such as learning, memory, andperception. On the clinical side, such a technology enables thedevelopment of closed-loop neurostimulation and neuroprosthetic systemsusing detailed neurophysiological signals for feedback. Beyond neuralrecording applications, long-lasting, high-fidelity neural interfacesdirectly benefit neurostimulation applications, as well.

Accordingly, this disclosure provides an innovative strategy that usesbiocompatible polymers to develop new micro-components that in certainvariations can be employed as electrodes in “microthread neural probes.”Micro-components formed in accordance with the principles of the presentdisclosure provide excellent electrical and transduction properties.Such micro-component-based microthread neural probes are ultra-small(e.g., having sub-cellular dimensions) and flexible, with bioactivesurfaces and nanostructured electrode sites for enhanced signaltransduction.

In certain aspects, such micro-component electrode probes can bereliably inserted into a brain and used for chronic recording, havingimproved performance in chronic neural recording of spike activity,which is considered to be the most sensitive assay of neural interfacematerial performance, for example. Design parameters of themicro-components include designing the size, flexibility, strength,conductivity, site electrical characteristics, insulating coating,insertion techniques, and electrode size.

Further, the inventive micro-components can be used in functionalizedmicrothread neural probes for targeted intervention in chronic tissueresponses. For example, immobilized biomolecules on a micro-component(e.g., microthread probe) surface can effectively interact with thesurrounding environment to elicit, augment, or minimize specificreactive tissue responses, including biofouling, inflammation, andneurotoxicity.

The implantable device is sized according to the size of the patient'sorgan in which it is to be implanted. The present disclosurecontemplates using micro-component devices independently implantedwithin target tissue or a target organ of a patient or alternately beingused in conjunction with or coupled to medical devices or other types ofmedical implants, known to those of skill in the art, which areintroduced and/or implanted internally in the patient. For example, tomonitor the brain, the neural probe can be directly implanted through aburr hole in the skull of the patient. By way of another non-limitingexample, such micro-electrodes can be used in cardiac pacemaker,monitoring assemblies, in/around peripheral nerves or a spine, under theskin, or in stents implanted in heart tissue or vasculature.

In various aspects, the inventive technology provides a micro-componentthat is relatively strong so that the micro-component is capable ofbeing incorporated into a device that can be implanted in vivo and isrelatively flexible to reduce potential stress at an interface withsurrounding tissue to mitigate cellular damage adjacent to themicro-component. While not limiting the present teachings to anyparticular theory, flexible micro-components incorporated intoimplantable devices (e.g., neural probes) appear to advantageouslyproduce less strain at a probe-tissue interface, and thus appear toadvantageously reduce tissue response in chronic implantations.Micro-components incorporated into implantable devices, for example, asmicroelectrodes, can be used for electrophysiological recordings as wellas recording the changes in concentration level of neural chemicals inthe brain (such as dopamine) or in the body (such as NO, glucose, pO₂).In other embodiments, devices incorporating such micro-components, forexample, as microelectrodes, may conduct electrical current or potentialfrom an external source, for example, as a probe or in a cardiacpace-maker application. Devices incorporating such micro-componentscreate long-lasting, high-fidelity neural interfaces, which mayoptionally further have biomimetic materials and surfaces. In certainvariations, such micro-components are incorporated into advancedimplantable neural probes for long-term (permanent), high quality andselective neural recording.

In various aspects, the micro-component is selected to have one or moredimensions that reduce damage in the surrounding tissue. As used herein,a “micro-component” has at least one spatial dimension that is less thanabout 100 μm (i.e., 100,000 nm), optionally less than about 50 μm (i.e.,50,000 nm), optionally less than about 10 μm (i.e., 10,000 nm), and incertain aspects less than or equal to about 5 μm (i.e., 5,000 nm).“Nano-sized” is generally understood by those of skill in the art tohave at least one spatial dimension that is less than about 50 μm (i.e.,50,000 nm), optionally less than about 10 μm (i.e., 10,000 nm),optionally less than about 1 μm (i.e., less than about 1,000 nm).

In various aspects, the dimensions of micro-component are of arelatively small scale, for example, on a microscale. A“micro-component” as used herein encompasses “nano-components.” Itshould be noted that so long as at least one dimension of themicro-component falls within the above-described micro-sized scale (forexample, diameter), one or more other axes may well exceed themicro-size (for example, length). In certain variations, amicro-component of the present teachings can comprise a micro-fiber,which has an evident longitudinal axis or axial geometry, and furtherhas at least one spatial dimension that is less than about 100 μm (i.e.,100,000 nm), optionally less than or equal to about 50 μm (i.e., 50,000nm). In certain preferred variations, a microfiber component has atleast one spatial dimension, such as a diameter, that is less than orequal to about 10 μm (i.e., 10,000 nm), optionally less than or equal toabout 9 μm (i.e., 9,000 nm), optionally less than or equal to about 8 μm(i.e., 8,000 nm), optionally less than or equal to about 7 μm (i.e.,7,000 nm), optionally less than or equal to about 6 μm (i.e., 6,000 nm),and in certain aspects less than or equal to about 5 μm (i.e., 5,000nm).

Thus, in certain aspects, depending upon the application, targetlocation, and individual patient, a micro-component of the presentteachings may have a major dimension, such as length, that is less thanor equal to about 150 mm (15 cm), optionally less than or equal to about100 mm (10 cm), optionally less than or equal to about 75 mm (7.5 cm),optionally less than or equal to about 50 mm (5 cm), optionally lessthan or equal to about 25 mm (2.5), optionally less than or equal toabout 10 mm (1 cm), optionally less than or equal to about 5 mm,optionally less than or equal to about 1 mm, optionally less than orequal to about 0.9 mm (900 μm), optionally less than or equal to about0.8 mm (800 μm), optionally less than or equal to about 0.7 mm (700 μm),optionally less than or equal to about 0.6 mm (600 μm), less than orequal to about 0.5 mm (500 μm), optionally less than or equal to about0.1 mm (100 μm), optionally less than or equal to about 75 μm,optionally less than or equal to about 50 μm, optionally less than orequal to about 40 μm, optionally less than or equal to about 30 μm,optionally less than or equal to about 25 μm, optionally less than orequal to about 20 μm, optionally less than or equal to about 15 μm, andin certain aspects, optionally less than or equal to about 10 μm. Incertain aspects, the micro-component has at least one major dimension(e.g., length) that is less than or equal to about 100 mm forimplantation into a human patient's brain as a neural probe.

Stated in another way, in certain aspects, an ultra-small implantabledevice incorporates a single micro-component having a cross-sectionalarea of less than or equal to about 70 μm², optionally less than orequal to about 65 μm², optionally less than or equal to about 60 μm²,optionally less than or equal to about 58 μm². In implantable devicescomprising an array of such micro-components, a cross-sectional area ofan ultra-small device can be less or equal to about 2,500 μm²,optionally less than or equal to about 2,000 μm², optionally less thanor equal to about 1,900 μm², optionally less than or equal to 1,800 μm²,for example.

In various aspects, the inventive technology provides a micro-componentthat comprises a conductive core material. In accordance with thepresent teachings, the core material has a high electricallyconductivity and therefore a low electrical resistivity, further isrelatively strong for implantation, and is selected to be relativelyflexible to reduce potential stress at an interface with surroundingtissue. In accordance with various aspects of the inventive technology,one or more discrete regions of a surface of the conductive corematerial are coated with an electrically conductive coating to formelectrically conductive surface regions, which can facilitate chargetransfer from the core material to an external conductor, such as anexternal electrode or other external lead. Such conductive regionsprovide nanostructured electrode sites for enhanced signal transductionin the micro-component.

The remaining regions of the surface of the core material, where theelectrically conductive coating is absent, are coated in accordance withthe present teachings with an electrically non-conductive (i.e.,electrically insulating) biocompatible coating to render these surfaceregions electrically insulated and non-conductive. In certainvariations, such an electrically non-conductive biocompatible materialcoating comprises a polymeric material that conforms to the surface ofthe core material to be insulated. In various aspects, themicro-component therefore comprises an electrically conductive corematerial having a surface with one or more discrete electricallyconductive regions, where the remainder of the exposed surface of thecore material has an electrically insulative coating. Such amicro-component can be employed as a micro-electrode for implantation asa medical device. While not limiting the inventive technology, inparticularly preferred embodiments the micro-component microelectrode isused as a neural probe implanted in vivo into an organism's brain.

In certain aspects, the conductive core material is optionally selectedto have a fiber shape. By “fiber” it is meant that the component definesan evident longitudinal axis and thus has a so-called “axial geometry.”Fibers having such an evident longitudinal axis include an elongatedaxial dimension, which is longer than the other dimensions (e.g.,diameter or width) of the fiber. In certain aspects, such elongatedfiber components having an axial geometry have an aspect ratio (AR)defined as a length of the longest axis divided by diameter of thecomponent, which is preferably at least about 100 and in certain aspectsgreater than about 1,000. In yet other aspects, such fibers may have anaspect ratio of 10,000 or more.

In various aspects, the core material has sufficient strength topenetrate into and through tissue, for example, brain tissue, withouttearing and with minimal buckling. In various aspects, it is preferredthat the electrically conductive core material comprises carbon. Incertain aspects, such an electrically conductive core may comprisegraphene. In certain embodiments, the core material comprises carbonnanotubes (CNTs), which are one-dimensional graphite sheets rolled up inthe shape of seamless, hollow cylinders. Single-walled carbon nanotubes(SWNT) are formed from a single sheet of graphite, while multi-walledcarbon nanotubes (MWNT) consist of multiple cylinders arranged in aconcentric fashion. The typical diameters of SWNT can range from about0.8 nm to about 2 nm, while MWNT can have diameters in excess of 100 nm.CNTs are known for their exceptional electrical conductivity, as well asmechanical properties. Metallic CNTs (e.g., CNTs exhibiting metallicbehavior) can carry electrical current density more than 1,000 timesgreater than metals such Au, Pt or Ir. Furthermore, CNTs haveexceptional tensile strength of σ=200 GPa and Young's modulus of Y≈1.2TPa as well as flexibility and robustness, large surface area, chemicalinertness and biocompatibility.

Carbon nanotube (CNT) films are strong, flexible, and conductive, with asufficient design space that support creation of ultrathin and flexiblemicrothread probes. The intrinsic properties of nanomaterials, such asnanocomposite CNT films, enable engineering the nano/micro organizationof the probe to meet physical, chemical and biological requirements.

Transferring the unique mechanical and electrical properties ofindividual nanoscale CNTs to macroscale composite polymer structures isa challenging area in CNT technology. CNT composites traditionallyprepared by polymerization and extrusion techniques have suffered fromshortcomings such as phase segregation especially at high CNTconcentration. The inability to load large amount of CNT into thecomposites have thus limited the ability to transfer the properties ofnanotubes to the matrix.

In accordance with the present teachings, a deposition technique can beused called layer-by-layer assembly (LBL) provides a reliable method forfabricating CNT-polymer composites with favorable characteristics. Theprinciple of the LBL technique relies on alternating adsorption ofpolyelectrolytes onto a substrate. The layers are built up by sequentialdipping of the substrate into oppositely charged polyelectrolytesolutions (FIG. 9, upper portion). Monolayers of individual componentsattracted to each other by electrostatic and van-der-Waals interactionsare sequentially adsorbed on the substrate.

Single-walled carbon nanotubes (SWNTs) are well suited for microthreadneural probes because of their unique mechanical, and electrical,properties. LBL films can be constructed on a variety of solidsubstrate, thus imparting much flexibility for size, geometry and shapeand further patterned or etched (with chemicals, plasma, electron beam,or high intensity lasers, for example). Following film fabrication bythe LBL process, the film can be detached from the substrate to yieldfree-standing, flexible and conductive samples. The electricalconductivity of the LBL films of SWNTs even in non-optimized conditionsis 1.5*10³ S/cm, with ultimate conductivity of individual SWNTs of 1*10⁴S/cm. This is at least one order of magnitude greater than theconductivity of the conducting polymer PEDOT, and equal or better thanthe conductivity of sputtered multicrystalline IrO_(x), which is oneorder less than that of the single crystals.

Additionally, LBL multilayers have both ionic and electronicconductivity that provides favorable charge transfer characteristics.The tubular morphology of the films is similar to that of PEDOT tubesthat show drastic increase of charge storage capacity and equallydrastic decrease of impedance.

In certain aspects, the tensile strength of SWNT LBL films is σ>430 MPa,which is substantially greater than that of polyimide filmsconventionally used as a backing for advanced implantable electrodes(PYRALIN, σ=350 MPa). Lateral orientation of nanotubes of LBL assembliesgreatly enhances their mechanical properties as opposed to the nanotube“forests” when CNTs are grown from the surface. Nanotubes attached tothe substrate only by the end can be easily broken off. This can occuras a result of tissue micromotion, as well as during the implantationprocedure.

A CNT film formed in accordance with these techniques is exceedinglyflexible in all directions, with excellent tensile strength. The thinpolymer electrically non-conductive coating, discussed in more detailbelow, is also flexible. When the microcomponent comprises a corematerial of CNT film, it can form a complete microthread that is(qualitatively) significantly more flexible than the best performingmicrofabricated thin-film polymer probes currently available, which aretypically 12-15 μm thick.

In certain embodiments, the core material comprises a carbon fiber,which has good flexibility, high electrical conductivity, and can bemade with a 5 to 10 μm diameter. As noted above, in various aspects, thecore material has sufficient strength to penetrate into and throughtissue, for example, brain tissue, without tearing and with minimalbuckling. In this regard, carbon fibers are particularly advantageous,since a typical 2 mm long intracortical implant carbon fiber has anominal value of stiffness (k)=(cross-sectional area×elasticmodulus)/length, 4,500 N/m and a nominal value of spring constant(k_(c))=3π(elastic modulus)[(outer diameter)⁴−(innerdiameter)⁴]/(64×length³) of about 0.01 N/m.

Certain desirable carbon fibers can have an elastic tensile modulus ofgreater than or equal to about 200 GPa, for example between 240 GPa toabout 999 GPa and can exhibit a modulus of elasticity up to 531 GPa,shear modulus of 2.2 GPa, and a tensile strength up to 5.65 GPa.Comparatively, the elastic modulus for typical carbon fibers (e.g., 234GPa) is far greater than that of silicon (164 GPa). Such physicalproperties make carbon fibers a desirable choice as a core material forthe micro-component electrodes in accordance with certain aspects of thepresent teachings. In other alternate variations, the conductive coremay comprise gold (which has a tensile modulus of 78.5 GPa), platinum,tungsten, steel, iridium, or a conductive composite material formed bylayer-by-layer assembly techniques. Thus, in certain aspects, theelectrically conductive core comprises a metal selected from the groupconsisting of gold, platinum, tungsten, steel, iridium, or combinationsthereof.

While not limiting the present disclosure to any particular theory, itis theorized that neural probes with a dimension under about ten, andmore desirably under about seven, micrometers reduce the foreign bodyresponse of an organism by preventing cellular adhesion or spreading.Therefore, in certain variations, the conductive core material comprisesa carbon fiber having a cross-sectional diameter of less than or equalto about 12 μm, optionally less than or equal to about 10 μm, optionallyless than or equal to about 9 μm, optionally less than or equal to about8 μm, in certain preferred aspects, optionally less than or equal toabout 7 μm, optionally less than or equal to about 6 μm, and in certainvariations, optionally less than or equal to about 5 μm.

The core material (e.g., CNT strand) is the electrically conductive partof the probe when it is incorporated into such a device as shown inFIGS. 8 and 12, with an exposed tip comprising the electricallyconductive region(s) or electrode site. A conformal insulative(non-conductive) polymer coating provides electrical insulation alongthe length of the core material, as well as a substrate forfunctionalization by attaching biomolecules. In certain embodiments, thecross-sectional dimension of a CNT strand is approximately 5×5 μm or 1μm×5 μm, which when combined with the approximately 0.5 to 1 μminsulator coating has desired sub-cellular size dimensions (thinnestassembled strand is about 2 μm).

In various aspects, the insulative electrically non-conductive materialis a polymer coating or optionally can also be a resistivelayer-by-layer coating, used both as dielectric coating with highimpedance and as reactive base layer for further surface modification.Reactive parylene coatings are particularly suitable, because they areconformal, thin, and therefore relatively flexible. Further, reactiveparylene coatings combine the well-established high resistivity ofconventional parylene with the ability to provide chemical anchor groupsfor subsequent modification with hydrogels or immobilization ofbiomolecules.

In various aspects, the insulative non-conductive material coating isbiocompatible. In various embodiments, biocompatible polymers are usedon a conductive carbon containing core, such as a carbon fiber, to makeultra-small neural probes that are flexible, yet durable, robust, andthat desirably have advanced bioactive capabilities for controllingintrinsic biological processes. In certain embodiments, the insulativematerial coating disposed on and forming the non-conductive regions ofthe core material's surface has a thickness of less than or equal toabout 1 μm, optionally less than or equal to about 950 nm, optionallyless than or equal to about 900 nm, optionally less than or equal toabout 850 nm, optionally less than or equal to about 800 nm.

The insulator coatings can be optionally functionalized. In variousembodiments, suitable insulator coatings can be vapor-deposited and/ormicro-patterned functional polymer coatings. In certain embodiments,suitable insulator coatings are based on a novel class ofvapor-deposited polymers referred to as functionalized poly-p-xylylenes(FIGS. 10 and 11). These polymer coatings are similar to thecommercially used parylene coatings in that they can be depositedconformally on a wide range of different substrates, but, in addition,provide versatile anchor groups for sophisticated surface modificationswith a wide range of different surface chemistries. Hence, thistechnology enables a one-step coating procedure to generatefunctionalized surfaces without requiring any kind of post-treatmentonce the films are deposited. The resulting polymers provide chemicallyreactive anchor groups for further surface modifications. Reactivecoatings are compatible with soft lithographic processes, allowing forpatterning of DNA, proteins, sugars, and mammalian cells, by way ofnon-limiting examples. In addition to the widespread availability of“anchor groups,” the simplicity in providing a wide range of functionalgroups, the excellent adhesion to various substrata, and itsapplicability to surfaces with three-dimensional geometries are keyadvantages when compared to polymers deposited by solvent-based methods.Further, such a technology has been successfully employed to developprotein resistant coatings.

Leading-edge reactive coatings can support regioselective immobilizationof bioligands via Husigen heterocycloaddition (so-called “clickchemistry”), or can result in highly non-fouling, protein- andcell-resistant surfaces. The latter is achieved by functionalCVD-coatings that can initiate atom-transfer radical polymerization(ATRP), followed by surface-induced graft-polymerization to createwell-defined hydrogel films. The use of CVD-based initiator coatingsprovides access to synthetic polymer hydrogels based on acrylates andmethacrylates, which have gained widespread popularity in variousbiomedical applications such as drug-delivery and tissue engineering.

Thus, in one embodiment, the insulative non-conductive coating comprisesa parylene-N insulator layer coated via chemical vapor deposition (CVD)on the surface of the core material, for example at an exemplarythickness of about 800 nm. In certain embodiments, the basicconfiguration of a single-strand microthread probe is a thin, flexibleand conductive CNT strand that is coated with a thin, conformalinsulative layer of functionalized poly-p-xylylenes (PPX—a biocompatiblecoating material closely related to parylene) along its length, exceptfor its tip (FIG. 8A). The core material (e.g., CNT strand) is theconductive part of the probe, with an exposed tip comprising theelectrically conductive region or electrode site. The conformalinsulative polymer coating provides electrical insulation along thelength, as well as a substrate for functionalization by attachingbiomolecules.

By way of non-limiting example, one insulative coating comprisesparylene-N and can be formed by one gram (1 g) of paracyclophane beingsublimed at 90-110° C. and 0.3 mbar and carried into the pyrolysischamber by argon at a flow rate of 20 sccm. After pyrolysis at 670° C.,the polymer is deposited on the substrate at 15° C. The deposition rate,according to the QCM, is 0.6-1.0 Å/s.

In certain aspects, the insulative non-conductive material coating istreated or comprises one or more biofunctional agents. By way of furtherbackground, neuronal loss and glial encapsulation surrounding neuralprostheses are well documented phenomena, although the mechanismsunderlying these changes in the device-tissue interface remain largelyunexplored. Inflammation in the organism plays a role; the release ofinflammatory cytokines from cells attached to explanted probes has beendemonstrated, and anti-inflammatory drugs such as dexamethasone havereduced astrogliosis associated with probes in vivo. Kim et al. reportevidence of reduced in vivo impedance associated withdexamethasone-releasing probes implanted in guinea pigs for a three weektime course (Kim and Martin 2006).

In recent years, a contribution of cell-cycle re-entry to the activationof glia, as well as neuronal apoptosis, has been demonstrated in modelsof central nervous system (CNS) injury. The cell cycle is a complexprocess through which cells progress from a quiescent to a proliferativestate, and cellular advancement through this cycle is controlled bycyclin dependent kinases (CDKs) and associated cyclins. Following CNSinjury and an associated upregulation of these cell-cycle proteins,mitotic cells (namely, microglia and astrocytes) proliferate, whilenon-mitotic cells (differentiated neurons) undergo caspase-mediatedapoptosis. Flavopiridol is a flavonoid drug which is a broad CDKinhibitor and arrests progression through the cell cycle. Di Giovanni etal. showed a single injection of flavopiridol intracerebroventricularlyreduced expression of the cell cycle protein cyclin D1, decreasedneuronal cell death, reduced glial activation, and improved motor andcognitive recovery following traumatic brain injury in rats.Flavopiridol has resulted in improved functional and histologicaloutcomes in vitro and in vivo models of spinal cord injury, Parkinson'sdisease, motor neuron apoptosis, and excitotoxic injury.

Since brain injury involves multiple biochemical pathways (i.e.,oxidative stress, excitotoxicity, and inflammation) and complexsignaling cascades, treatment strategies that target multiple mechanismsappear to be an optimal solution. In other medical pathologies, such ascancer and AIDS, the effect of multiple therapeutic agents has proven tobe greater than that of the sum of the single counterparts.Advantageously, the micro-component electrodes for implantable neuralmedical devices prepared in accordance with the present disclosure havethe ability to present multiple biofunctional or therapeutic agents tothe surrounding environment and tissue (e.g., combinational drugtherapy), which is believed to provide benefits for therapy fortraumatic brain injury. By way of non-limiting example, flavopiridol andminocycline have been observed to have synergistic effects in preservinghippocampal neurons following ischemic brain injury in rats. Alternativedosing regimens are an additional opportunity for optimizing effects onrecording stability and quality.

In various aspects, the micro-component electrodes of the presentdisclosure provide one or more of the following attributes that aredesirable for enhanced long-term tissue interface with an implantedmedical device: (1) Minimizing biofouling—the adsorption of non-specificproteins to probe surfaces—including for example, blood-borne proteinsresulting from disruption of the blood-brain barrier during probeinsertion—which can provide inflammatory signals and degrade electricalcharacteristics; (2) Minimizing immune responses involving cellularresponses (primarily astrocytes and microglia) that tend to encapsulateprobes over time; and (3) Minimizing neurotoxic processes involving bothdirect neuron damage, as well as deleterious extracellular signaling,oxidative stress, and the like.

Therefore, in certain embodiments, immobilized biomolecules on theexposed microcomponent surfaces can be effective for intervening withspecific reactive tissue responses, including biofouling, inflammation,and neurotoxicity. It is believed that micro-component surfaces that donot become bio-fouled (i.e., have little non-specific proteinadsorption) elicit subsequent attenuated reactive tissue responses andimproved electrical characteristics. Therefore, in certain embodiments,the electrically non-conductive biocompatible material coating istreated or comprises one or more agents that i) reduce the accumulationand/or attachment of microorganisms or cells to the surface of theinsulative coating, therefore serving to minimize or reduce bio-foulingor other undesirable cellular growth and/or ii) otherwise interact withthe surrounding environment in a predetermined way. “Biofilm resistant”or “biofouling resistant” refers to any coating or agent that impairs,inhibits, prevents or retards the attachment and/or growth of biofoulingorganisms or cells.

In certain aspects, the insulative electrically non-conductive coatingmay comprise one or more “biofunctional” or “bioactive” substances,which refers to a chemical substance, such as a small molecule, activeingredient, macromolecule, metal ion, or the like, that causes anobservable change in the structure, function, optical function, orcomposition of a cell when a cell is exposed to such a substance. Invarious aspects, the biofunctional substance or agent promotes cellregeneration, differentiation, growth, proliferation, and/or repair, forexample.

Functional bio-coatings and anti-biofouling coatings, such aspoly(ethylene glycol) (PEG), are beneficial for their ability to improvechronic neural interfaces. Therefore, in certain variations, the presentdisclosure provides methods of making a micro-component electrode thatcomprises coating an electrically conductive core material with anon-conductive material. The non-conductive material can further befunctionalized. For example, in one embodiment, the micro-component corematerial surface is coated with an 800 nm layer of parylene-N viachemical vapor deposition and then modified by using a functionalizedparylene with poly(ethylene glycol) (PEG) poly(ethylene glycol)-terminalmethacrylate (PEGMA) as a biofunctional substance. Poly(ethylene glycol)methacrylate can be immobilized onto the parylene surface by severaldifferent methods, including but not limited to atom transfer radicalpolymerization and grafting to a functionalized coating.

By way of example,poly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)] isoptionally used as initiator for atom transfer radical polymerization(ATRP) because of its functional groups. After chemical vapor depositionof poly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)], themicro-component electrodes are incubated under inert conditions withdegassed aqueous solution of oligo(ethylene glycol) methyl ethermethacrylate, with CuBr/CuBr2/bpy as the catalyst. The polymerizationsproceed at room temperature for 4 hours. Microthread electrodes arethoroughly rinsed after the reaction.

In one exemplary method used to create the initiator coatings for ATRPvia CVD polymerization,poly(p-xylylene-4-methyl-2-bromoisobutyrate-co-p-xylylene) (2) in FIG.10 is prepared via CVD polymerization in a custom-made CVDpolymerization system. The starting material,[2.2]paracyclophane-4-methyl 2-bromoisobutyrate (1), is sublimed undervacuum and converts by pyrolysis into the corresponding quinodimethanes,which spontaneously polymerizes upon condensation to the cooledsubstrate surface, which is maintained at 15° C. Throughout CVDpolymerization, a constant argon flow of 20 sccm and a working pressureof 0.5 mbar are maintained. The pyrolysis temperature is set to be 550°C. and sublimation temperatures are between 115-125° C. under theseconditions. CVD polymerization spontaneously occurs on samples placed ona rotating, cooled sample holder. In cases, where patterned substratesare required, polydimethylsiloxane (PDMS) micro-stencils can be sealedto the substrates during CVD polymerization. Next, an aqueous solutionof the methacrylate, 2,2′-bipyridine (bpy), and CuBr₂ is stirred in aSchlenk flask at room temperature. The homogeneous solution can bedegassed using three freeze-pump-thaw cycles. Next, copper bromide isadded under nitrogen purge to the frozen solution and the molar ratio ofCuBr/CuBr₂/bpy can be set to be 1/0.3/2.5. The mixture is warmed up toroom temperature and stirred until a homogeneous dark brown solution isformed. The solution is then transferred into a nitrogen-purgedglovebag, and the polymerizations proceed at room temperature for a setreaction time. Samples are analyzed in triplicate. To prepare samplesfor protein adsorption and cell adhesion studies, the CuBr concentrationis 10 mM and polymerizations is allowed to proceed for three hours atroom temperature.

The use of CVD-based initiator coatings provides access to a variety ofbiofunctional materials, including synthetic polymer hydrogels based onpolyethylene glycol methacrylates (PEGMA), which have protein resistantproperties. Beyond PEGMA, certain methacrylate-based monomers withdifferent side functional groups are contemplated, which impartprotein-resistant properties to the coated surfaces. These monomers canbe deposited after CVD polymerization via graft-co-polymerization (FIGS.10 and 11) from ATRP initiator groups. In certain aspects, monomers are2-hydroxyethyl methacrylate (HEMA), poly(ethylene glycol) methyl ethermethacrylate (PEGMA), 3-sulfopropyl methacrylate potassium salt (SMPS),[2(methacryloyloxy)ethyl]dimethyl-(3-sulfopropyl) ammonium hydroxide(MEDSAH), and combinations thereof.

In certain variations, during ATRP polymerization, anti-inflammatorymolecules, such as dexamethasone can be added to the polymerizationsolution and loaded into the hydrogel layer. In this way, a depot layeris created that can release dexamethasone during initial implantperiods, thereby reducing the early onset of inflammatory responses.Alternatively, interlukin (IL-10) may be used as anti-inflammatorycomponent. IL-10 is a small protein and is therefore structurally andfunctionally different from dexamethasone. Dexamethasone and IL-10 canbe incorporated together during this process, as well. A variety ofother biofunctional materials can be incorporated into the insulatorcoating in this manner.

In another exemplary method for grafting PEG to the parylene coating,poly(p-xylylene carboxylic acid pentafluorophenolester-co-p-xylylene) isselected because it is a polymer with an active ester group that canreadily react with primary amine. After chemical vapor deposition ofpoly(p-xylylene carboxylic acid pentafluorophenolester-co-p-xylylene),micro-component electrodes are incubated in 10 mM mPEG-NH₂ (MW 10,000)PBS solution for 8 hours, followed by thorough rinses.

FIGS. 11A-11F shows an exemplary formation of a parylene coating that isfurther functionalized by ATRP applied to a conductive core material toform a micro-component of the present disclosure. Such a micro-componenthas a sub-cellular cross-sectional dimension, but is flexible, stronger,and with sufficient electrical characteristics for neural recording andadvanced bioactive capabilities for controlling intrinsic biologicalprocesses. 7 μm diameter (tensile modulus 234 GPa) carbon fibers arefirst mounted onto a microelectrode printed circuit board. The carbonfiber is then coated with an 800 nm poly(p-xylylene) coating viachemical vapor deposition (CVD) polymerization (FIG. 11A).Poly(p-xylylene) (also know under the trade name Parylene-N™) isselected for its very low dissipation factor, high dielectric strength,and low dielectric constant that is also invariant with frequency.

One, two, or three individual 7 μm diameter (tensile modulus=234 GPa)carbon fibers are mounted onto a NeuroNexus A16 printed circuit board ora bare stainless steel wire using silver epoxy (WPI; Sarasota, Fla.) andbaked at 140° C. for 10 min. An approximately 800 nm thickpoly(p-xylylene) insulator layer is then coated via CVD. One gram ofparacyclophane is sublimed at 90-110° C. and 0.3 mbar and carried intothe pyrolysis chamber by argon at a flow rate of 20 standard cubiccentimeters per minute (sccm). After pyrolysis at 670° C., the polymeris deposited on the substrate at 15° C. The deposition rate, accordingto the QCM is 0.6-1.0° A/s.

An additional 50 nm thick layer of the functionalized polymer coatingpoly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)] isdeposited onto the device via CVD polymerization (FIG. 11B). Thispolymer provides initiator groups for subsequent atom transfer radicalpolymerization (ATRP). After ATRP, a poly(ethylene glycol methacrylate)(PEGMA) top layer is formed that is about 200 nm thick (FIG. 11C), whichrenders the neuronal probe devices protein-resistant.

Poly(ethylene glycol methacrylate) is grafted onto the poly(p-xylylene)surface by atom transfer radical polymerization (ATRP).Poly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)] is usedas initiator for ATRP because of its functional groups. After CVDpolymerization ofpoly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)], MTEs areincubated under inert conditions with a degassed aqueous solution ofoligo(ethylene glycol) methyl ether methacrylate, with CuBr/CuBr2/bpy asthe catalyst. The polymerizations proceed at room temperature for 4hours. MTEs are thoroughly rinsed after the reaction.

In certain aspects, the methods of making a micro-component electrodefurther comprises coating the electrically conductive core material withan electrically conductive material, to form a recording site, forexample. Such a process may be preceded by cutting a portion of theconductive core material having the applied non-conductive coating andthen applying the electrically conductive material to the exposedregion. For example, a recording site is created by electrochemicaldeposition of poly(3,4-ethylenedioxythiophene) poly(styrenesulfonate)(PEDOT/PSS) onto the tip of the neuronal probe, from which thepoly(p-xylylene) and poly(ethyleneglycol) methacrylate coatings areremoved (FIG. 11D). Poly(p-xylylene)-coated and PEGMA-grafted carbonfibers are cut to a length of 0.3-0.5 cm. For PEDOT deposition, monomer3,4-ethylenedioxythiophene (EDOT) (Bayer, Germany) is electrochemicallypolymerized and deposited onto the surface of the electrode sitestogether with the anions in the solution. Specifically, PEDOT/PSS iselectropolymerized from a 0.1 M poly(sodium-p-styrenesulfonate) (PSS)(Acros Organics; Morris Plains, N.J.) aqueous solution with an EDOTconcentration of 0.01 M under galvanostatic conditions. In galvanostaticmode, the current is held at 100 pA. FIG. 11D shows removal ofinsulation by cutting away the tip to expose a carbon site, which isthen electrodeposited with PEDOT.

The combination of such advanced materials provides an ultra-smallorganic interface that has the approximate size of a single trace of aconventional silicon neural probe, but advantageously has sufficientstrength and flexibility to act as a stand-alone independent electrode.

FIGS. 12A-12D show exemplary schematics and cross section ofarchitecture of a neural probe incorporating a micro-component electrodein accordance with one aspect of the present teachings. FIG. 12A shows afirst side view of an exemplary neural probe on a printed circuit board(10-8 mm) and FIG. 12C shows a second exemplary neural probe on aprinted circuit board (5-3 mm) made in a manner similar to thatdescribed above in the context of FIGS. 11A-11F. In the sectional view(FIG. 12C of the probe in FIG. 12A), a conductive core comprises carbon,which is coated with a parylene coating, followed by a PPX-PEG coating.FIG. 12D shows a detailed sectional view of the layers forming anexemplary MTE probe, where the conductive core comprises carbon, whichis coated with a parylene coating, followed by a PPX-PEG coating. At aterminal end, the probe is cut and PEDOT is applied.

In certain aspects, the CVD technology provides bioconjugation of activebiomolecules to the micro-component electrode surface. While thebioconjugation chemistry can be easily adapted by changing the couplinggroups incorporated in the CVD polymer, an exemplary group discussedherein is neural cell adhesion molecule (NCAM). NCAM can exhibitneuroprotective as well as neurotrophic properties. Another exemplarycoupling group for immobilization along the insulator coating surface isthe growth brain-derived neurotrophic factor (BDNF) growth factor, usedindependently or in conjunction with NCAM. Notably, various proteins,ligands, saccharides, and other bioactive molecules can be selected andused in various combinations for immobilization of proteins onto CVDfilms. In certain aspects, the orientation of biofunctional substances(requiring specific active regions to be exposed to the surroundingenvironment) on the surface of the insulator coating can be achieved bya variety of methods. Such techniques include using a reaction of BDNFwith Msc-ONSu, so that protective groups are successively introduced ina protein yielding pure fractions isolated by ion exchangechromatography. Glycosaminoglycans, such as heparin, are alsoimmobilized onto CVD coatings. Simultaneous immobilization of twodifferent proteins, here NCAM and BDNF, in controlled ratios usingbio-orthogonal immobilization strategies are contemplated.Quantification of surface-bound biomolecules is done by surface plasmonresonance spectroscopy (Biacore X) on CVD-coated gold slides.

In various aspects, the micro-components have one or more conductiveregions formed on the surface of the conductive core material, to serveas nanostructured electrode sites for enhanced signal transduction tothe core material from an external electrical source. Such electricallyconductive electrode sites optionally have geometric surface areasranging from about 100 μm² to 500 μm² (corresponding to an uninsulatedtip about 5 to 25 microns long assuming all sides are not coated), whichis consistent with sizes of conventional microelectrodes for unitrecording.

In yet other aspects, at least one of the electrically conductive corematerial or the electrically non-conductive biocompatible coating of theimplantable micro-component electrode are optionally patterned, shaped,and/or etched. Thus, the electrically conductive core material or theelectrically non-conductive biocompatible coating of the implantablemicro-component electrode can be treated prior to implantation. Suchpatterning, shaping, and/or etching can be achieved by techniques thatemploy wax or resists, photolithography, electron beam, laser, and/orplasma treatment, by way of non-limiting example.

In certain embodiments, one or more of the electrically conductiveregions may be formed by removal of the insulative electricallynon-conductive coating after application, for example, by mechanicalremoval (e.g., cutting and/or scraping) or by ablation (e.g., laserablation). In other embodiments, one or more of such regions may beformed by masking the desired areas along the surface of the corematerial prior to application of the electrically non-conductiveinsulative coating. Using a wax coating for example, the tip of a carbonfiber can be shielded from parylene coating during the CVD process.Then, after applying the parylene electrically non-conductive coatingover the wax, the wax can be melted away in hot ethanol. In otheraspects, such regions may be formed by removing the insulative coatingand then applying the electrically conductive material to the regions.By way of non-limiting example, a carbon electrode site (approximately38.5 μm²) is exposed by cutting a discrete region of the surface of theinsulator-coated (parylene-coated) carbon core material to create theregion to be coated with a conductive material.

One particularly suitable electrically conductive material is abiocompatible conductive polymer, poly(3,4-ethylene dioxythiophene)(PEDOT) with a poly(4-styrenesulfonate) (PSS) counter-ions. Such aPEDOT:PSS conductive polymer is electrochemically deposited onto theexposed carbon core material to decrease recording site impedance. ForPEDOT deposition, monomer 3,4-ethylenedioxythiophene (PEDOT) (Bayer) iselectrochemically polymerized and deposited onto the surface of theelectrode sites together with the anions in the solution. Specifically,PEDOT/PSS is electropolymerized from a 0.1 Mpoly(sodium-p-styrenesulfonate) (PSS) (Acros Organics; Morris Plains,N.J.) aqueous solution with a PEDOT concentration of 0.01 M under eithergalvanostatic conditions. In galvanostatic mode, the current is variedfrom 50 to 500 pA. Presence of PEDOT/PSS is confirmed with cyclicvoltammetry and electrochemical impedance spectroscopy. Other suitableconductive biocompatible materials for coating the one or moreconductive regions on the surface of the core material includepolypyrrole, platinum, platinum black, iridium oxide, carbon nanotubes,graphene, and combinations thereof.

In various aspects, an impedance of a core material comprising carbon istypically about 2.5 MΩ to about 6 MΩ. Thus, in various embodiments, theconductive materials for the one or more conductive regions on thesurface are selected to reduce the impedance of the core materialsurface at the interface electrically communicating with an externalelectrical lead. For example, conductive coatings comprising PEDOT canhave an impedance of less than or equal to about 500 kΩ to greater thanor equal to about 10 kΩ. In certain aspects, the PEDOT may have animpedance ranging from about 10 to about 500 kΩ; and in certain aspects,may range from greater than or equal to about 100 kΩ to less than orequal to about 500 kΩ. Other suitable materials include platinum black(PtB), which typically has an impedance of greater than or equal toabout 800 kΩ to less than or equal to about 2.5 MΩ.

In certain aspects, the present disclosure incorporates themicro-components into a single-strand microthread probe (having theelectrically conductive region/site at the tip) and 1-dimensional linearelectrode arrays (array of conductive sites positioned along the axialdimension). In certain embodiments, the multi-strand probes optionallyhave multiple conductive region sites (for example, up to 8 sites, with30 to 200 μm site separations). These types of base assemblies can bemodified to more complex two- and three-dimensional microthread probearrays, as well.

In certain variations the present teachings contemplate theincorporation of the micro-components into assemblies that formmicrothread arrays for long-term and high fidelity in vivo (e.g., neuralinterfaces). While not limiting the present disclosure in any particularmanner, microelectrode probes of the present teachings are capable ofthe following: (1) recording useful spike activity from cortex anddeeper structures for a duration of more than six months; (2) a superiorrecording quality; (3) increased stability and diminished degradation ofin vivo recordings. Therefore, in certain variations the micro-componentelectrodes of the present disclosure increase the quality, stability,and longevity of neural recordings.

In certain embodiments, microthread electrodes (MTEs) comprisingmicro-components are prepared by mounting individual 7 μm diameter(tensile modulus 234 GPa) carbon fibers onto an acute microelectrodeprinted circuit board. (FIG. 8). The device is then coated with 800 nmparylene-N coating via chemical vapor deposition CVD. A carbon electrodesite (approximately 38.5 μm²) is exposed by cutting the parylene coatedcarbon. See for example FIGS. 1A-C. Contact pins are gently scrapedusing a scalpel, and a carbon electrode site is exposed by cutting theparylene coated carbon fiber with a pair of scissors. Poly(3,4-ethylenedioxythiophene) (PEDOT) with a poly(4-styrenesulfonate) (PSS)counter-ion is electrochemically deposited onto the exposed carbon toprovide an electrically conductive region on the surface of the carboncore that serves as a recording site and to decrease recording siteimpedance (FIGS. 1C-E). Such a neural implant has ultra-small dimensions(e.g., 38.5 μm² electrode size and 58 μm² footprint), while alsodemonstrating flexible, robustness, and durability.

Cyclic voltammetry (CV) measurements are made on parylene coated fibers,parylene coated fibers with a approximately 38.5 μm² exposed carbon tip,and a parylene coated fiber with an exposed PEDOT/PSS recording site(FIG. 2A). CV profiles indicate the presence of PEDOT on the recordingsite. Impedance spectroscopy (EIS) measurements displayed progressivedecrease in impedance (FIGS. 2B-2D) with increased PEDOT deposition.Also as expected, charge storage capacity increased with longer PEDOTdeposition durations (FIG. 2E).

The microthread probes are manually inserted into rat cerebral cortex tovalidate the insertion technique. In vivo cortical recordings with PEDOTdeposited, parylene insulated carbon fibers are able to record neuralspikes in rat motor cortex. (FIGS. 3A-B). In all in vivo trials, thePEDOT coated MTE is able to detect at least one neuronal spike with asignal to peak-to-peak-noise ratio (SNR) of greater than 1.1 ranging upto 8.0. In contrast, comparative parylene insulated carbon fibers with acut carbon exposed recording site are unable to record any neuronalspikes with an SNR greater than 1.1. The PEDOT deposited MTEs is able torecord local field potential (LFP) activity. A comparative uninsulatedcarbon fiber lacking the conductive PEDOT coating implanted 2 mm intothe cortex is able to record LFPs, but unable to discriminate any singleunit spikes—indicating the necessity of insulating the fiber to providea more localized recording environment.

Thus, comparative parylene insulated carbon fibers with a cut carbonexposed recording site (without a conductive coating) is unable torecord any neuronal spike with an SNR greater than 1.1. Interestingly,the PEDOT MTE prepared in accordance with the present disclosure hadhigher peak to peak noise then the carbon site MTE (FIGS. 3C-3E). Theincreased noise floor may be due to biological noise. The PEDOT MTE isable to pick up neuronal spikes from many distant neurons to increasethe noise floor. By using a carbon site MTE reference for example, moreneurons can be discriminated from the background.

Spectrograms show more detailed low-frequency activity as a function oftime during a typical recording segment (FIGS. 4A-C). Settings for thespectrogram are the same as with the power spectral densities, exceptsliding 1-second windows are used to observe changes throughout a shortrecording session. Recordings from the PEDOT fiber depict oscillatoryactivity around 4-Hz in the theta band that occurs in burst throughoutthe recording session (FIG. 4A); this is typically referred to as aketamine spindle bursting activity. Recordings from the bare carbon siteare dominated by a 60-Hz instrumentation noise throughout the session,which is not observed from the PEDOT fiber. However, low frequencytheta-band activity is sufficiently large in amplitude to be observedfrom the carbon electrode, even though individual neural spike activityis not present during the recording sessions.

In summary, novel microelectrodes are provided with reduced feature sizethat is able to record single unit spikes. This is the smallest neuronalrecording electrode site that PEDOT has been grown onto. It is also thesmallest implanted probe to successfully record single unit activity inthe cortex. The stiffness (k) of an exemplary MTE prepared in accordancewith the present teachings is calculated from the cross sectional area,elastic modulus and length and 2 mm length is 4,500 N/m which is anorder of magnitude smaller than a commercially available siliconmicroelectrode of the same length, 151,000-246,000 N/m. The springconstant (k_(c)) of an exemplary MTE is calculated to be 0.01 N/m whilea commercially available silicon microelectrode of the same length is2.13-3.46 N/m in the planar dimension and 149-615 N/m in the lateraldimension.

FIG. 4A depicts the power spectral densities of the neural recordingsshowing the intensity of the recordings as a function of frequency.These are created using a Hamming window for smoothing with a32768-point FFT. Low frequency activity in the range typically observedfor local field potentials is presented. The peak at 0-Hz is indicativeof a slight DC-offset in the signal in the recordings. Peaks observed at4-Hz and 10-Hz are representative of low-frequency synchronous activityin the theta and alpha bands respectively. Pronounced theta-bandactivity is typically observed under ketamine anesthesia, which is usedhere.

In certain alternate embodiments, the one or more conductive regions onthe core material (e.g., the recoding sites) are made from other highimpedance materials such as gold (gold can be further modified—forexample by using self-assembled monolayers). A wax is placed onto asilicon wafer and spun to a certain thickness. The length of the exposedcarbon site can be controlled by controlling the thickness of the wax onthe wafer, where faster spins leads to a thinner wax layer. CVD ofparylene is carried out. Then, the parylene over the wax and the wax canbe melted away in hot ethanol. In such embodiments, the micro-componentcan be used as an electrode for chemical sensing, such as dopamine.

In certain aspects, the present disclosure provides methods ofmonitoring, sensing, or stimulating neural activity in an organism. Inone aspect, such a method comprises electrically communicating with aprobe implanted into an organism. In certain variations, the methodcomprises electrically communicating with a neural probe implanted intoa brain of an organism. The probe optionally has a cross-sectional areaof less than or equal to about 2,000 micrometers-squared (μm²) andcomprises at least one micro-component electrode. The at least onemicro-component electrode comprises an electrically conductive corematerial having a surface with one or more electrically conductiveregions disposed on the surface of the electrically conductive corematerial. The one or more electrically conductive regions disposed onthe surface of the electrically conductive core comprise an electricallyconductive polymeric coating. Further, an electrically non-conductivepolymeric coating is disposed on regions of the surface of theelectrically conductive core material corresponding to locations wherethe one or more electrically conductive regions are absent. In certainpreferred variations, the micro-component electrode has at least onedimension less than or equal to about 10 μm.

In certain variations of the present technology, a neural probe is ableto detect at least one neuronal spike with a signal to peak-to-peaknoise ratio (SNR) of greater than or equal to about 1.1 or a signal totwo-standard deviation noise ratio (SNR) of greater than or equal toabout 2. Further, the one or more electrically conductive regions maycomprise platinum, platinum black, carbon nanotube, carbon, orcombinations thereof. Thus, the one or more electrically conductiveregions can be used to detect neural chemicals, biochemicals, and/orother chemical agents.

In yet other aspects, the methods of the present teachings may includeintroducing the neural probe into the organism (e.g., implanting viasurgical techniques) prior to the methods where the neural probe is usedto electrically communicate with the target tissue in the organism. Suchintroducing can include at least one of the following implantationprocesses: pulling the neural probe, sewing the neural probe into,accelerating or injecting the neural probe, wrapping the neural probearound, or otherwise disposing and placing the neural probe adjacent toa target tissue of the organism. The introducing may include implantingthe neural probe into a brain, or in certain other variations, maycomprise implanting the neural probe near a spine, a peripheral nerve,vasculature, or an organ of the organism.

For example, a removable shuttle and/or a stiff dissolvable coating(such as crystallized PEG) may be used on the implantable probe(micro-component electrode) to assist with or facilitate the insertionof the micro-thread probe or an array of micro-thread probes into theorganism. For example, a shuttle may be coated with a hydrophilic layerto assist in the separation of the shuttle and the probe after insertionto aid in the removal of the shuttle. Alternatively, a probe may beinserted into the tissue with high velocity (e.g., by injecting theprobe with force). Further, in another aspect, it may be pulled or sewninto the target tissue (with a larger conventional shank or a needle) orwrapped around a nerve.

EXAMPLE 1

One, two, or three individual 7 μm diameter (Tensile Modulus 234 GPa)carbon fibers are mounted onto a NeuroNexus A16 print circuit boardusing silver epoxy (WPI; Sarasota, Fla.) and baked at 140° C. for 10minutes. An approximately 800 nm thick parylene-N insulator layer isthen coated via CVD, in a deposition chamber.

A reaction system for CVD polymerization comprises a sublimation zone, apyrolysis zone, and a deposition zone. The CVD installation shouldprovide (i) flexible control of polymerization parameters, (ii)monitoring capability to measure critical polymerization parameters insitu, and (iii) on line feedback capability for instant process control.CVD polymerization typically involves base pressures of around of 5×10⁻⁵bar and working pressures (i.e. the pressure during polymerization) of0.05 to 0.4 mbar. The pressure control unit is located at thedown-stream part of the equipment while the carrier gas inlet and flowcontrol is located up-stream. This allows for controlling workingpressure within a flexible range of gas flows and also prevents pressurefluctuations due to sublimation of the precursor.

One gram of paracyclophane is sublimed at 90-110° C. and 0.3 mbar andcarried into the pyrolysis chamber by argon at a flow rate of 20 sccm.After pyrolysis at 670° C., the polymer is deposited on the substrate at15° C. The deposition rate, according to the QCM, is 0.6-1.0 Å/s.

Parylene coated carbon fibers are then cut to a length of 0.3-0.5 cm.For PEDOT deposition, monomer 3,4-ethylenedioxythiophene (EDOT) (Bayer)is electrochemically polymerized and deposited onto the surface of theelectrode sites together with the anions in the solution. Specifically,PEDOT/PSS is electropolymerized from a 0.1 Mpoly(sodium-p-styrenesulfonate) (PSS) (Acros Organics; Morris Plains,N.J.) aqueous solution with an EDOT concentration of 0.01 M undergalvanostatic conditions. In galvanostatic mode, the current is variedfrom 50 to 500 pA.

Poly-ethylene glycol is immobilized onto the parylene surface by twodifferent methods.Poly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)] is usedas initiator for atom transfer radical polymerization (ATRP) because ofits functional groups. After chemical vapor deposition ofpoly[(p-xylylene-4-methyl-2-bromoisobutyrate)-co-(p-xylylene)],microthread electrodes are incubated under inert conditions withdegassed aqueous solution of oligo(ethylene glycol) methyl ethermethacrylate, with CuBr/CuBr2/bpy as the catalyst. The polymerizationsproceeded at room temperature for 4 hours. Microthread electrodes arethoroughly rinsed after the reaction.

In another technique, poly(p-xylylene carboxylic acidpentafluorophenolester-co-p-xylylene) is a polymer with an active estergroup which can readily react with primary amine that is grafted to thefunctionalized coating. After chemical vapor deposition ofpoly(p-xylylene carboxylic acid pentafluorophenolester-co-p-xylylene),microthread electrodes are incubated in 10 mM mPEG-NH₂ (MW 10,000) PBSsolution for 8 hours, followed by thorough rinses.

SEM imaging is carried out on a FEI Quanta 200 3D Focussed Ion BeamWorkstation. Samples are sputter coated with gold prior to imaging.

The microthread probes are manually inserted into rat cerebral cortex tovalidate the insertion technique in a typical experimental preparation.Adult male Sprague-Dawley rats (Charles River Laboratories) weighing300-350 g are prepared for cortical implants using previouslyestablished methods (Ludwig et al., 2006; Vetter et al., 2004). Theanimal is anesthetized with a mixture of 50 mg/mL ketamine and 5 mg/mLxylazine administered intraperitoneally with an initial dosage of 0.125mL/100 g of body weight and regular updates of ketamine. The depth ofanesthesia is regulated by monitoring heart rate and blood oxygensaturation. The animal is placed into a stereotaxic frame and a 2 mm by2 mm craniotomy is made over the motor cortex. The dura is incised andresected. Sterile saline is used to keep the brain surface moistthroughout the procedure. A stereotaxic frame mounted micromanipulatorguided insertions of the microthread electrode 2 mm into the cortex. Dueto the stiffness of the carbon fiber, a 2-5 mm long fiber did not needadditional assistance in inserting into the cortex when insertedperpendicularly to the cortical surface using a stereotaxic guide.

In another aspect, a removable shuttle and/or a stiff dissolvablecoating (such as crystallized PEG) may be used to assist in theintroduction (e.g., insertion) of the microthread probe or an array ofmicrothread probes into the organism. The shuttle may be coated with ahydrophilic layer to assist in the separation of the shuttle and theprobe after insertion to aid in the removal of the shuttle.Alternatively, it may be inserted into the tissue with high velocity.Further, in another aspect, it may be pulled or sewn into the tissue(with a larger conventional shank or a needle) or wrapped around anerve.

For in vivo neural recordings in the laboratory animals,electrophysiological data are acquired using a TDT RX5 Pentusa RecordingSystem (Tucker-Davis Technologies, Alachua, Fla.). These neuronalsignals are acquired through a head-stage buffer amplifier to avoidsignal loss in data transmission. Signals are sequentially filtered byan antialiasing filter in the preamplifier, digitized at an approximate25-kHz sampling rate, and digitally band-pass filtered from 2 to 5000Hz. Wideband signals are acquired to capture both spiking and LFPactivity. Signals are continuously recorded in segments ranging from 30seconds to >10 minutes in duration.

Neural recording segments are analyzed offline to determine the numberof neurons recorded, noise levels, and signal amplitudes using customautomated MatLab (Mathworks Inc., MA) software, as described in detail84 and utilized elsewhere. As an overview, the wide-band recordings arefiltered in software to isolate the spike data (300-5000 Hz) from theLFP data (1-100 Hz). To identify individual units, the threshold for thehigh-frequency data is established by using a window set at 3.5 standarddeviations below the mean of the data. A 2.4-msec waveform is extractedfrom the data stream at each threshold crossing. To group isolatedwaveforms to a single neuronal unit, principal component analysis isthen completed, and the resultant components are separated intoindividual clusters by using Fuzzy C-means clustering. Units withsufficiently clustered principal components are plotted, and thesignal-to noise ratio is calculated as the peak-to-peak amplitude of themean waveform of the cluster divided by six times the standard deviationof the remaining data stream after all waveforms had been removed.

Impedance spectroscopy measurements (EIS and CV measurements) are madeusing an Autolab potentiostat PGSTAT12 (Eco Chemie, Utrecht, theNetherlands) with associated frequency response analyzer and generalpurpose electrochemical system software (Brinkmann, Westbury, N.Y.),respectively. To obtain EIS and CV measurements, each probe is submergedin a phosphate buffered saline (PBS) solution of 137 mM sodium chloride,2.7 mM potassium chloride, and 11.9 mM phosphate buffer with a stainlesssteel rod serving as the counter electrode and a standard Ag/AgCl probeas reference. Impedance measurements are taken between 10 Hz and 31 kHzat 25 mVrms. CV values are obtained by cycling three times from 0.8V to−0.6V at a sweep rate of 1 V/s and averaging the data. Charge storagecapacity (CSC) of each site is calculated from the full area under theCV curve, scaled by the inverse of the scan rate. After implantation, adistant stainless steel (316-SS grade) bone screw is used as thereference and counter electrode.

In certain aspects, a microfabricated, thin polymer structure preparedin accordance with the principles of the present teachings can beattached to a larger, conventional shank such that the electrode sitecan be placed on the edge (FIGS. 5A-B and 6B). The micro-components ofthe present disclosure can be incorporated into parylene-substrateneural probes having a stiff penetrating shank (48 μm by 68 μm)supporting a thin lateral extension (5-μm thick and 100-μm wide). Probestructures fabricated with a sub-cellular dimension significantlyreduced encapsulation while preserving healthy neurons around the thinedge (FIG. 7).

The most dramatic result is a significant decrease in encapsulationdensity around the lateral edge (L) relative to the shank (S).Encapsulation density at the lateral platform edge is almost ⅓^(rd) thelevel around the device shank (129% and 425% increase in first 25 μm,respectively, P=0.00003, N=7 animals) (FIG. 7C). Significant decrease inneuronal loss is also observed. Neuronal loss is reduced from 30% to 48%in the first 25 μm from the interface (FIG. 3D). Qualitatively, a numberof important results are also found. The ramified morphology of themicroglia (O×42) indicated less activation, and astrocyte (GFAP)structure indicated only moderate hypertrophy (FIG. 7A). Mostencouraging, microglia and other nonneuronal cells did not conform tothe edge of the structure. The thin parylene structure (5 μm thick)elicited greatly reduced cellular encapsulation with no distinctcapsular boundary.

Layer-by-Layer assembly of carbon nanotubes are used to form a corematerial of the inventive micro-components in certain variations of thepresent disclosure. Thus the core material optionally comprises ananotube component made by a layer-by-layer assembly process employing asingle-walled carbon nanotubes (SWNT) obtained from CarbonNanotechnologies Inc. Since the LBL process relies on sequentialadsorption of the oppositely charged components, the SWNT isnon-covalently modified by the adsorption of a negatively chargedpolymer, poly(sodium-4-styrenesulfonate) (PSS, Sigma-Aldrich). The SWNTthin film is fabricated on a solid substrate such as silicon wafers orsilicon probe mockups by sequentially dipping the substrate intodispersions of SWNT-PSS and polymer solutions. The individual layeringstep, adsorption of SWNT and polyelectrolyte monolayers, is interceptedby a rinsing and drying step to remove any excess solution and preventcross contamination. The dipping time at each step of the layeringprocess is varied to achieve best quality films but is not expected toexceed 20 min. Films of different thicknesses (or layers) are produced.Thick films (>100 layers) are used throughout this study for cellular,mechanical, electrical and electrochemical investigations. The progressof the film assembly is monitored by a variety of techniques, includingUV-VIS, ellipsometry, atomic force microscopy (AFM) and scanningelectron microscopy (SEM). Following LBL assembly, free standing filmsare achieved by detaching the SWNT films from the solid substrate bytreatment with hydrofluoric acid (HF), a common etching solution. Theas-prepared film on substrate is gently loaded into a bath of 5% aqueousHF solution until detachment of the film from the solid substrate isobserved (few minutes). The detached film is rinsed in water and buffersolution to remove residual HF. Such a film is then optionally used acore material of the inventive micro-components.

EXAMPLE 2

In certain aspects, the present disclosure pertains to variations of themicro-component comprising platinum black coatings. By way ofbackground, carbon fiber ultramicroelectrodes are known for in vivodopamine sensing (as well as ascorbic acid and pH sensing) throughamperometry (constant potential or cyclic voltammetry), and have alsobeen used for electrophysiology recordings and enzymatic biosensing.Their small size makes them particularly attractive as a platformrecording neural signals. In accordance with the present disclosure,carbon fibers have electrodeposited platinum black functionalizationlayers. In certain aspects, platinum black coatings offer severalimportant advantages. First, platinum black increases the functionalsurface area, while having little to no change on the geometric area.Second, the platinum black coatings can be used to modulate electrodeimpedance, which can improve electrophysiology recordings. Third,platinum black can increase the charge capacity of the electrode forstimulation. Finally, as compared to bare carbon, platinum black hashigher catalytic activity for the oxidation of hydrogen peroxide, whichis a common detector molecule of enzymatic biosensors for choline,glutamate, glucose, and other neurochemicals. Carbon conductive regions(of the carbon fibers) as well as carbon nanotube or graphene-depositedconductive regions can also be used for chemical sensing.

An 800 nm insulation layer of parylene-n is deposited on a 7 μm diametercarbon fiber via chemical vapor deposition. Parylene-n is chosen as analternative to the standard glass insulation to maintain flexibility ofthe fiber electrode. After parylene-n deposition, the end of the fiberis cut to expose a bare carbon tip (38.5 μm²). Platinum black iselectrochemically deposited on the exposed tip. The deposition ischaracterized and confirmed by impedance, cyclic voltammetry, andscanning electron microscopy. Using constant potential amperometry (700mV vs. Ag/AgCl), robust sensitivity to hydrogen peroxide is demonstratedhaving applicability to peroxide-based biosensing mechanisms. Initialacute in vivo implantation in rat motor cortex showed the ability torecord neuronal spikes with high fidelity. Such flexible carbon fiberelectrodes having platinum black layers in the conductive surfaceregions provide a modular platform for neural recordings.

Details of insertion, such as tip shape, speed, and insertion locationare important considerations for chronic neural interfaces. There isincreasing evidence that the details of electrode insertion, such aselectrode shape and insertion speed, may impact tissue damage. Detailsof probe insertion may also impact chemical trauma in tissue, as theimplantation of microelectrodes punctures and tears neural vasculature.During insertion, the highly regulated blood brain barrier iscompromised leading to plasma protein release into the surroundingparenchyma, resulting in adsorption onto the electrode, increasedconcentrations of extracellular serum proteins, ions, and other solutes,and deposition of plasma into the neuropi. Acutely, this damage can beobserved as irregularities in neuronal spike activity, elevated levelsof extracellular neurotransmitters and a net increase in extracellularwater content. This vasogenic edema leads to increased brain tissuevolume and intracranial pressure often associated with their impact ontissue damage and clinical outcome. This initial acute damage leads tothe release of erythrocytes, clotting factors, and inflammatory factorsfrom disrupted blood vessels, which facilitate recruitment of activatedmicroglia and a broad region of astrocyte activation around the insertedprobe. For example, albumin, the most abundant protein in blood plasma,is responsible, in part, for inducing glial cell activation andproliferation. Although stab wound studies show limited chronic tissuedamage, plasma protein adsorption onto the electrode may perpetuate thetissue response in chronically implanted electrodes. Whileanti-biofouling coatings have issues with stability when implantedoutside of the central nervous system (CNS), in certain aspects, it istheorized that the coatings on implantable devices in the CNS may onlyhave to be stable and effective long enough for the cells to clean upthe plasma exudates.

Furthermore, disrupting major arterioles during probe insertion cancause additional neuronal damage, by means of ischemia, through loss ofperfusion to tissue below the disrupted region, which is typically wherethe recording sites are located. Drastically reducing feature size,stiffness, and applying a self-assembled monolayer anti-biofoulingcoating as provided by various aspects of the present inventivetechnology will likely reduce these acute insertion trauma effects.

Chronic in vivo electrophysiology is demonstrated in FIGS. 14A-14L.Initial chronic neural recording studies suggest that the biosensorsprepared in accordance with certain aspects of the present teachings arestable over several weeks in the brain. For example, exemplary MTEsdemonstrated high yield across animals for detecting single unitactivity overtime (FIG. 14A—showing a percentage of active chronicallyimplanted MTEs able to detect at least 1 single unit (dashed line) as afunction of weeks post-implant (n=7)), which is comparable or betterthan the yield of conventional silicon devices or wire microelectrodearrays. The mean SNR of the largest discernable single unit detected byeach MTE shows the stability of the chronic single unit recordings (FIG.14B). During the initial inflammation period, the SNR decreases, thenincreases temporarily to a peak around one week before stabilizing.Tissue response to stab wounds subsides by four weeks suggestingstabilization of the tissue after surgery. For this reason,electrophysiological recordings are continued up to five weeks. By thesecond week the single units appear to stabilize as indicated by thedecreased fluctuation and standard deviation. This is stabilizationperiod is similar to previously reported studies. Deconstructing themean SNR of the largest discernable single unit to the units' amplitudeand the electrodes' noise floor shows that the fluctuations in SNR inthe first two weeks is due to fluctuations largely in signal amplitudeas opposed to fluctuations in noise floor (FIG. 14C). Examiningindividual electrodes for SNR or signal amplitude fluctuations furtherdemonstrate the large variability observed in the first week afterimplantation, and stabilization after the second week (FIG. 14D). FIGS.14E-14F demonstrate two examples of over 5 weeks of chronic recordingsfrom two different animals. These initial chronic recording studiesdemonstrate that the inventive biosensors remain intact in the brainover moderate lengths of time, without any electrophysiological evidenceof signal degradation. They also have extremely high yield of unitrecordings compared to conventional Silicon-on-Anything (SOA) probes.

The foregoing description of the embodiments has been provided forpurposes of illustration and description. It is not intended to beexhaustive or to limit the invention. Individual elements or features ofa particular embodiment are generally not limited to that particularembodiment, but, where applicable, are interchangeable and can be usedin a selected embodiment, even if not specifically shown or described.The same may also be varied in many ways. Such variations are not to beregarded as a departure from the invention, and all such modificationsare intended to be included within the scope of the invention.

What is claimed is:
 1. An implantable micro-component electrodecomprising: a core consisting of an electrically conductive corematerial; one or more electrically conductive regions disposed directlyon the surface of the electrically conductive core material comprisingan electrically conductive biocompatible coating; an electricallynon-conductive biocompatible coating disposed directly on regions of thesurface of the electrically conductive core material where the one ormore electrically conductive regions are absent; and a layer comprisinga biofunctional agent disposed directly on the electricallynon-conductive biocompatible coating, the biofunctional agent beingpoly(ethylene glycol) (PEG) or poly(ethylene glycol methacrylate)(PEGMA), wherein the micro-component electrode has at least onedimension of less than or equal to about 10 μm.
 2. An implantablemicro-component electrode of claim 1, wherein the electricallyconductive core material that has an elastic tensile modulus of greaterthan or equal to about 200 GPa and a nominal value of stiffness ofgreater than or equal to about 5 GN/μm and the electricallynon-conductive biocompatible coating has a thickness of less than 1 μm.3. The implantable micro-component electrode of claim 1, wherein theelectrically conductive core comprises carbon.
 4. The implantablemicro-component electrode of claim 3, wherein the electricallyconductive core is selected from the group of materials of carbonfibers, composites comprising single-walled carbon nanotubes, compositescomprising multi-walled carbon nanotubes, graphene, and combinationsthereof.
 5. The implantable micro-component electrode of claim 1,wherein the electrically conductive core comprises a metal selected fromthe group consisting of: gold, platinum, tungsten, steel, iridium, orcombinations thereof.
 6. The implantable micro-component electrode ofclaim 1, wherein the electrically conductive coating comprises aconductive biocompatible material selected from the group consisting of:poly(3,4-ethylene dioxythiophene) (PEDOT) with apoly(4-styrenesulfonate) (PSS) counter ions, polypyrrole, platinum,iridium oxide, carbon nanotubes, graphene, and combinations thereof. 7.The implantable micro-component electrode of claim 1, wherein theelectrically non-conductive biocompatible coating comprises a parylenepolymer or a parylene polymer derivative.
 8. The implantablemicro-component electrode of claim 1, wherein at least one of theelectrically conductive core material or the electrically non-conductivebiocompatible coating are patterned, shaped, and/or etched.
 9. Animplantable neural probe medical device comprising one or more of themicro-component electrode of claim
 1. 10. A micro-component electrodefor an implantable medical device comprising: a core consisting of anelectrically conductive core material comprising carbon and having adiameter of less than or equal to about 10 μm; one or more electricallyconductive regions disposed directly on the surface of the electricallyconductive core material comprising an electrically conductive polymericcoating; and an electrically non-conductive polymeric coating disposeddirectly on regions of the surface of the electrically conductive corematerial where the one or more electrically conductive regions areabsent, wherein the electrically non-conductive coating comprises aparylene polymer or a parylene polymer derivative.
 11. Themicro-component electrode of claim 10, wherein the electricallyconductive core is selected from the group of materials of carbonfibers, composites comprising single-walled carbon nanotubes, compositescomprising multi-walled carbon nanotubes, graphene, and combinationsthereof.
 12. The micro-component electrode of claim 10, wherein theelectrically conductive polymeric coating comprises a conductivebiocompatible material selected from the group consisting of:poly(3,4-ethylene dioxythiophene) (PEDOT) with apoly(4-styrenesulfonate) (PSS) counter ion, polypyrrole, carbonnanotubes, graphene, and combinations thereof.
 13. The micro-componentelectrode of claim 10, wherein the electrically non-conductivebiocompatible coating further comprises one or more biofunctionalagents.
 14. An implantable neural probe medical device comprising one ormore of the micro-component electrodes of claim
 10. 15. A method ofmonitoring, sensing, or stimulating neural activity in an organism, themethod comprising: electrically communicating with a neural probeimplanted into a brain of the organism, the neural probe having across-sectional area of less than or equal to about 2,000 μm² andcomprising at least one micro-component electrode comprising a coreconsisting of an electrically conductive core material having a surfacewith one or more electrically conductive regions disposed directly onthe surface of the electrically conductive core material comprising anelectrically conductive polymeric coating; an electricallynon-conductive polymeric coating disposed directly on regions of thesurface of the electrically conductive core material corresponding tolocations where the one or more electrically conductive regions areabsent; and a layer comprising a biofunctional agent disposed directlyon the electrically non-conductive biocompatible coating, thebiofunctional agent being poly(ethylene glycol) (PEG) or poly(ethyleneglycol methacrylate) (PEGMA), wherein the micro-component electrode hasat least one dimension less than or equal to about 10 μm.
 16. The methodof claim 15, wherein the neural probe is able to detect at least oneneuronal spike with a signal to peak-to-peak noise ratio (SNR) ofgreater than or equal to about 1.1 or a signal to two-standard deviationnoise ratio (SNR) of greater than or equal to about
 2. 17. The method ofclaim 15, further comprising introducing the neural probe into theorganism prior to said electrically communicating.
 18. The method ofclaim 17, wherein the introducing comprises implanting the neural probenear a spine, a peripheral nerve, vasculature, or an organ of theorganism.
 19. The implantable micro-component electrode of claim 1,wherein the layer comprising a biofunctional agent comprises an activebiomolecule selected from the group consisting of neural cell adhesionmolecule (NCAM), brain-derived neurotrophic factor (BDNF),glycosaminoglycans, and combinations thereof.
 20. The implantablemicro-component electrode of claim 13, wherein the one or morebiofunctional agents is at least one of poly(ethylene glycol) (PEG) andpoly(ethylene glycol methacrylate) (PEGMA).
 21. The implantablemicro-component electrode of claim 10, wherein the electricallynon-conductive polymeric coating further comprises an active biomoleculeselected from the group consisting of neural cell adhesion molecule(NCAM), brain-derived neurotrophic factor (BDNF), glycosaminoglycans,and combinations thereof.
 22. The implantable micro-component electrodeof claim 1, wherein the electrically conductive core material is acarbon fiber.